One of the central goals of tissue engineering and related fields is to create large, whole organ replacements containing a physiologically relevant blood vessel network for the treatment of human disease or injury. While progress has been made in the past several decades, clinical successes have been limited to avascularized constructs for the replacement of thin tissues such as skin (Yannas et al., 1982; Heimbach et al., 1988), cornea (Nishida et al., 2004), and bladder (Atala et al., 2006). The remaining major barrier to creating large, whole organ replacements is one of nutrient and waste transport (Ko et al., 2007; Johnson et al., 2007; Khademhosseini and Langer, 2007). Diffusion alone is sufficient for the growth of human cell aggregates up to several hundred micrometers thick. However, cells at the center of larger cell clusters cannot adequately access nutrients or remove waste by diffusion due to obstruction and metabolism from adjacent cells and accordingly, large cell aggregates develop necrotic cores. What is needed for the successful engineering of living organs or in vitro models of solid tumor growth and malignancy is a combination of convective and diffusive transport—a vascular network.
To enable convective transport within biocompatible materials, three approaches are currently being investigated. The most common approach utilizes various material processing steps such as critical point drying (Dagalakis et al., 1980), gas foaming and salt leaching (Jun and West, 2005), or electrospinning (Pham et al., 2006) to create macroporous structures that can be perfused in vitro for tissue culture. Unfortunately, most of these additional processing steps require reagents or conditions that are cytotoxic and cannot be done in the presence of living cells. Furthermore, because these materials have an open, porous void volume rather than a vessel network, they cannot be matched up with native vasculature in vivo to permit blood flow (Ko et al., 2007). A second approach utilizes co-cultures of endothelial cells and smooth muscle progenitor cells to create random capillary-like structures in 3D inside biomaterials such as collagen gels (Koike et al., 2004). However, the several day time delay required for these cells to generate their own vasculature means that cells at the center of constructs even a few hundred micrometers in diameter may die from lack of adequate nutrient and waste transport before a vessel network is formed. The third approach utilizes photolithographic equipment from the microprocessor industry to create microfluidic structures in a layer-by-layer fashion (Ling et al., 2007; Golden and Tien, 2007). Photolithography requires expensive, proprietary equipment to reach micrometer-scale resolution, and lithography is typically done in cartesian coordinates to yield channels with rectangular cross-sections. In contrast, native vasculature rarely follows straight x-, y-, or z-vectors, and blood vessels have circular cross-sections, meaning that microfluidic scaffolds may not resemble native vasculature sufficiently to recapitulate organ function. Moreover, this mode of fabrication is much too slow to make large models of organ vasculature efficiently. For example, to fabricate only a 1 cm3 model organ structure containing vasculature that makes up 10% of its total volume with micrometer-scale resolution by stereolithography would require patterning 100 billion individual voxels in a serial fashion; even allowing only 1.6 microseconds per voxel (Hahn et al., 2006) this would take 44 hours. At this rate, it would take 3.8 years of continuous fabrication time to create a single construct with the 750 mL of vasculature found in an adult human liver. Extreme optimization or less stringent resolution requirements may allow dramatic time savings, but the exorbitant cost of the equipment and technical expertise needed for this process foreshadows considerable financial hurdles for mass production. While layer-by-layer photomasking is already much faster than raster scanning because a single xy plane of voxels can be fabricated in parallel (Liu Tsang et al., 2007), difficulties in aligning successive layers with high precision and edge-to-edge artifacts that are typically found in these structures highlight the significant technical challenges that are generally associated with photolithography for 3D fabrication at micron-scale resolution. A similar approach of direct-writing in 3D involves a custom-built polyelectrolyte liquid ink extruder which can print three-dimensional webs of microperiodic structures (Gratson et al., 2004; Therriault et al., 2005). However, the resulting vessels were uniform in diameter, required long timescales for deposition, and had ill-defined inlet and outlet geometry which did not resemble physiologic vasculature.
Basic anatomy demonstrates that identical organs from different people have vascular architectures unique to each of them; yet these organs can still function similarly for each person. Thus, it is not necessarily the exact x, y, and z coordinates of individual vessel components that allow proper functioning for an organ. Rather, the overall transport of blood components that results from vessel architecture is the principal factor what defines healthy and diseased tissue (e.g. vessel tortuosity, red blood cell velocity, pO2, and pH; see FIG. 1).